In commercially available X-ray diagnostic devices, equipment for X-ray imaging, the components such as for example X-ray tube, X-ray image detector, etc. can move relative to one another during the image recording by lateral shift or displacement distances smaller than the pixel size, without thereby significantly affecting the image information and image resolution. These innocuous displacements created by movement lie in the range of ten to a few hundred micrometers.
In contrast, displacements of the magnitude of the grating period, which typically measure only a few micrometers, produce a dramatic deterioration in the image information in a Talbot-Lau apparatus. This results in considerable demands on mechanical stability if a conventional X-ray system is to be upgraded into a Talbot-Lau apparatus which uses the method described above to determine the image information.
For a C-arm X-ray recording device there is the particular difficulty that the X-ray tube and X-ray detector are located at the ends of a C-arm, permitting vibrational motions which can result in lateral mechanical displacements of the two components of approx. a hundred micrometers.
In the Talbot-Lau apparatus for X-ray phase contrast imaging, as described for example by Georg Pelzer et al. in “Energy-resolved interferometric X-ray imaging”, published in Proc. SPIE 8668, Medical Imaging 2013: Physics of Medical Imaging, (Mar. 19, 2013) pages 866851 ff., the gratings G0, G1 and G2 are mounted on a fixed rack. By means of a controlled lateral displacement of one of these gratings G0 to G2 in fractions of the grating period, for example by a piezo-stepper, an intensity distribution is scanned (phase stepping), from which the image variables absorption, phase displacement and dark field can be calculated.
An X-ray recording system, with which differential phase contrast imaging of the type mentioned in the introduction can be performed, is known for example from U.S. Pat. No. 7,500,784 B2, which is explained on the basis of FIG. 1.
FIG. 1 shows the typical essential features of an X-ray recording system for an interventional suite with a C-arm 2 held by a stand 1 in the form of a six-axis industrial or articulated robot, with an X-ray source, for example an X-ray emitter 3 with X-ray tube and collimator, and an X-ray image detector 4 being attached to the ends of said C-arm 2 as an image recording unit.
By way of the articulated robot known for example from U.S. Pat. No. 7,500,784 B2, which preferably has six axes of rotation and thus six degrees of freedom, the C-arm 2 can be adjusted spatially as required, being rotated for example about a center of rotation between the X-ray emitter 3 and the X-ray detector 4. The inventive angiographic X-ray system 1 to 4 is rotatable in particular about centers of rotation and axes of rotation on the C-arm plane of the X-ray image detector 4, preferably about the center point of the X-ray image detector 4 and about axes of rotation intersecting the center point of the X-ray image detector 4.
The known articulated robot has a baseframe which for example is permanently mounted on a base. To this a carousel is rotatably attached about a first axis of rotation. Attached to the carousel so as to pivot about a second axis of rotation is a robot rocker arm, to which is fixed a robot arm which can rotate about a third axis of rotation. A robot hand is attached at the end of the robot arm, so as to rotate about a fourth axis of rotation. The robot hand has a fixing element for the C-arm 2, which can pivot about a fifth axis of rotation and can rotate about a sixth axis of rotation extending perpendicular thereto.
The implementation of the X-ray diagnostic device is not dependent on the industrial robot. Normal C-arm devices can also be used.
The X-ray image detector 4 can be a rectangular or square, flat semiconductor detector that is preferably made of amorphous silicon (a-Si). Integrating and possibly counting CMOS detectors can also be used, however.
Located in the beam path of the X-ray emitter 3 on a tabletop 5 of a patient positioning couch is a patient 6 to be examined, as an examination object. On the X-ray diagnostic device a system control unit 7 is connected to a high-voltage generator for generating the tube voltage and to an image system 8 which receives and processes the image signals from the X-ray image detector 4 (operating elements are not shown, for example). The X-ray images can then be viewed on displays of a monitor bracket 10 held by means of a ceiling-mounted support system 9 which can travel lengthwise, pivot and rotate, and is height-adjustable. Also provided in the system control unit 7 is a processing circuit 11, the function of which is further described below.
Instead of the X-ray system illustrated by way of example in FIG. 1 with the stand 1 in the form of the six-axis industrial or articulated robot, the angiographic X-ray system can, as shown in a simplified manner in FIG. 2, also have a normal ceiling- or floor-mounted bracket for the C-arm 2.
Instead of the C-arm 2 shown by way of example, the angiographic X-ray system can also have separate ceiling- and/or floor-mounted brackets for the X-ray emitter 3 and the X-ray image detector 4, which for example are electronically fixedly coupled.
In the arrangements currently being focused on for clinical phase contrast imaging, conventional X-ray tubes, currently available X-ray image detectors, as described for example by Martin Spahn [2] in “Flat detectors and their clinical applications”, Eur Radiol, Vol. 15, pages 1934 to 1947, and three gratings G0, G1 and G2 are used, as explained in greater detail below on the basis of FIG. 2, which shows a schematic structure of a Talbot-Lau interferometer for differential phase contrast imaging with expanded tube focus, gratings G0, G1 and G2, and a pixelated X-ray image detector 4.
To generate coherent radiation the X-ray beam 13 exiting from a tube focus 12 of the non-coherent X-ray emitter 3 penetrates an absorption grating 14 (G0), which produces spatial coherence of the X-ray source, as well as an examination object 15, for example the patient 6.
Because of the examination object 15 the wave front of the X-ray beam 13, which is represented by its normal 16, is deflected by phase displacement, as is made clear by the normal 17 of the wave front without phase displacement, i.e. without an object, and the normal 18 of the wave front with phase displacement. Then the phase-displaced wave front passes through a diffraction or phase grating 19 (G1) with a grating constant adjusted to the typical energy of the X-ray spectrum for generating interference lines or interference patterns 20, and in turn an absorbent analyzer grating 21 (G2) for reading out and capturing the interference pattern 20 generated. The grating constant of the analyzer grating 21 is adjusted to that of the phase grating 19 and of the remaining geometry of the Talbot-Lau arrangement. The analyzer grating 21 is arranged e.g. at the first or n-th Talbot distance. The analyzer grating 21 here converts the interference pattern 20 into an intensity pattern, which can be measured by the X-ray image detector 4. Typical grating constants for clinical applications are in the order of a few μm, as can also be inferred for example from the cited reference [1].
If the tube focus 12 of the beam source is sufficiently small and the radiation output generated is nevertheless sufficiently large, it may be possible to dispense with the first grating G0, the absorption grating 14, as is the case if for example a plurality of field emission X-ray sources is provided as an X-ray emitter 3, as is known from DE 10 2010 018 715 A1 which is described below.
The differential phase displacement is now determined for each pixel element of the X-ray image detector 4 in that using so-called “phase stepping” 22, which is indicated by an arrow, the analyzer grating 21 G2 is displaced in several steps by a corresponding fraction of the grating constant perpendicular to the normal 16 of the wave front of the X-ray beam 13 or 17 and laterally to the arrangement of the grating structure, and the signal Sk arising for this configuration during the recording is measured in the pixel of the X-ray image detector 4 and thus the resultant interference pattern 20 is scanned. For each pixel element the parameters of a function (e.g. sine function) describing the modulation are determined by a suitable fit procedure, an adjustment or equalization procedure, to the thus measured signals Sk. The visibility, i.e. the standardized difference between maximum and minimum signal, is here a measurement for characterizing the quality of a Talbot-Lau interferometer. It is defined as a contrast of the scanned modulation
  V  =                              I                      ma            ⁢                                                  ⁢            x                          -                  I                      m            ⁢                                                  ⁢            i            ⁢                                                  ⁢            n                                                I                      ma            ⁢                                                  ⁢            x                          +                  I                      m            ⁢                                                  ⁢            i            ⁢                                                  ⁢            n                                =          A              I        _            
In this equation A further designates the amplitude and Ī the average intensity. The visibility can assume values between zero and one, since all variables are positive and Imax>Imin. In a real interferometer it is also the case that Imin>0, so that the value range of V is expediently exploited. Minimum intensities greater than zero and all non-ideal properties and shortcomings of the interferometer result in a reduction in the visibility. A third item of information, which can be defined by way of the visibility and is generated by this mode of measurement, is designated as the dark field. The dark field indicates the ratio between the visibilities of measurement with an object and those without an object.
  D  =                    V        obj                    V        ref              =                            A          obj                ·                              I            _                    ref                                      A          ref                ·                              I            _                    obj                    
Three different images can then be generated from the comparison of particular derived variables from the fitted functions for each pixel, once with and once without an object (or patient):
(i) absorption image,
(ii) differential phase contrast image (DPC) and
(iii) dark field image.